The present embodiments relate to an activation method for a number of field coils of a magnetic coil system of a magnetic resonance system.
Magnetic resonance systems include, for example, a basic magnet that generates a temporally static, spatially essentially homogeneous basic magnetic field in an examination volume. The basic magnetic field may have a strength greater than 1 tesla. Basic magnetic field strengths of 1.5 tesla or 3 tesla may be used. An examination object (e.g., a patient) is introduced into the examination volume. The examination object is excited by a high-frequency system to emit magnetic resonance signals that are received by the same high-frequency system or a different high-frequency system and subsequently evaluated.
For spatial encoding, additional magnetic fields are overlaid on the basic magnetic field during excitation of the magnetic resonance signals, between excitation and the receiving of the magnetic resonance signals, and during the receiving of the magnetic resonance signals by field coils of a magnetic coil system of the magnetic resonance system. The overlaid magnetic fields may be oriented in the same direction as the basic magnetic field but are spatially different. The directions, along which the additional magnetic fields differ, may define a rectangular Cartesian coordinate system. The field coils may bring about an essentially linear field change (e.g., gradient fields) within the examination volume.
The magnetic fields generated by the field coils are followed very exactly with respect to the spatial field distribution. Otherwise, artifacts result during image reconstruction (e.g., when evaluating the received magnetic resonance signals). For example, the field coils are produced with a high level of mechanical precision. The field coils of the magnetic coil system may, however, only be produced with a mechanical precision in the millimeter range. It is difficult to achieve higher accuracy due to the tolerances of the parts used and the structural tolerances (e.g., a large number of individual layers have to be disposed on the smallest space). If high-voltage resistance and mechanical strength are to be achieved for field strengths above 1 tesla, there are further tolerances due to the vacuum casting process. Also, with the actively shielded field coils, the radial distance between the field-generating primary layer and the shielding layer is of significance. This distance is also subject to certain tolerances.
The residual scatter fields that are inevitable due to the finite length of the field coils and the finite number of conductor loops together with the scatter field due to tolerances induce an eddy current in the conductive surfaces of the magnetic resonance system. These induced eddy currents distort the spatial and temporal pattern of the magnetic fields generated in the examination volume and are therefore a potential cause of image artifacts.
The zero and first order components of the interference fields may be compensated for with the aid of the known eddy current compensation (ECC) method. The remaining higher-order interference fields vary with the manufacturing tolerance of the field coils and may not be compensated for in a conventional magnetic coil system. In the case of a gradient coil system, interference above the first order in the region of approximately 0.1% of a desired gradient strength is treated as acceptable.
In the case of magnetic resonance imaging, diffusion-weighted magnetic resonance sequences are used. To mark the magnetization (e.g., diffusion encoding), diffusion-weighted magnetic resonance sequences use a number of strong gradient pulses with a high amplitude-time integral. The orientation of the pulses defines the diffusion direction to be investigated in each instance. Diffusion sequences of the current prior art (e.g., DSI, HARDI, q-ball) measure proton diffusion not only in six orthogonal orientations but in many more orientations (e.g., 60 orientations). This is associated with a measurement time in the region of 10 minutes to 1 hour. The information thus obtained is used by tractography procedures to predict the pattern of nerve pathways (e.g., in the brain). Conventional gradient systems allow an amplitude of approximately 40 mT/m for such measurements. In the case of new systems that have a number of field coils for each gradient direction, this value increases to up to 300 mT/m. The amplitude of the eddy currents excited by the strong gradient pulses increases to the same degree.
To reduce interference fields, an improved production method may be used for the field coils (e.g., manufacturing the field coils with a higher level of precision). Once the field coils have been produced, the coil sensitivities may be fine-tuned. A further known solution is the dynamic activation (ECC) of additional higher-order field coils. This last solution is, however, associated with a high additional outlay, since at least one additional field coil is required for each interference field type. Also, for physical reasons, not all the higher-order interference fields may be compensated for independently of one another.
An activation method of the type mentioned in the introduction is known from DE 199 55 117 A1. With this method, the activation signals of the activation signal vector are determined using an optimization calculation. A target function that is a function of the activation signals and a measure of the deviation of the ideal field from the target field is established. The target function is minimized.